An exemplary analyte sensor 10 with which aspects of the invention may be employed is shown in FIG. 1. The sensor may be implantable subcutaneously or in the intraperitoneal fluid and is configured for detecting an analyte of interest, such as glucose. The sensor 10 includes a sensor body 12, a matrix layer (or indicator) 14 coated over part or all of the exterior surface of the sensor body 12, with fluorescent indicator molecules 16 distributed throughout the layer, a light source 18, e.g. an LED, that emits light, including light over a range of wavelengths which interact with the indicator molecules (referred to herein simply as “light at a wavelength which interacts with the indicator molecules”), i.e., in the case of a fluorescence-based sensor, a wavelength which causes the indicator molecules 16 to fluoresce, and a photosensitive element 20, e.g. a photodetector, which, in the case of a fluorescence-based sensor, is sensitive to fluorescent light emitted by the indicator molecules 16 such that a signal is generated in response thereto that is indicative of the level of fluorescence of the indicator molecules.
In some embodiments the indicator molecules are contained within the matrix layer 14, which comprises a biocompatible polymer matrix that is prepared according to methods known in the art and coated on the surface of the sensor body. Suitable biocompatible matrix materials, which must be permeable to the analyte, include hydrogels which, advantageously, can be made selectively permeable—particularly to the analyte—i.e., they perform a molecular weight cut-off function. The sensor 10 may also include reflective coatings 32 formed on the ends of the sensor body 12, between the exterior surface of the sensor body and the matrix layer 14, to maximize or enhance the internal reflection of the light and/or light emitted by fluorescent indicator molecules.
The sensor 10 may be configured to be implantable in a patient and may be constructed in such a way that no electrical leads extend into or out of the sensor body to supply power to the sensor (e.g., for driving the source 18) or to transmit signals from the sensor. Thus, the sensor may include a means for receiving power from an external source 40 that is wholly embedded or encapsulated within the sensor body 12 and a transmitter 42 that also is entirely embedded or encapsulated within the sensor body 12.
As still further illustrated in FIG. 1, an optical filter 34 may be provided on the light-sensitive surface of the photosensitive element (photodetector) 20. This filter prevents or substantially reduces the amount of light generated by the source 18 from impinging on the photosensitive surface of the photosensitive element 20. At the same time, the filter allows fluorescent light emitted by fluorescent indicator molecules 16 to pass through it to strike the photosensitive region of the detector. In addition, a temperature sensor 64 and an optional signal amplifier 66 may also be provided. The temperature sensor 64 measures the locally surrounding temperature of the ambient tissues and the indicator molecule environment and provides this information to the control logic circuit (not shown). The control logic circuit correlates fluorescence level, for example, with analyte concentration level, thereby correcting the output signal for variations affected by temperature. Amplifier 66 is a relatively simple gain circuit which amplifies the signal generated by the photodetector 20.
The various components and circuitry of the sensor 10 may be assembled onto a ceramic (e.g., ferrite) substrate 70.
Exemplary sensors are described in U.S. Pat. Nos. 5,517,313; 6,330,464; and 6,400,974, as well as in U.S. patent application Ser. No. 13/761,839, the respective disclosures of which are hereby incorporated by reference in their entireties.
If a sensor is implanted in the body of a living animal, the animal's immune system may begin to attack the sensor. For instance, if a sensor is implanted in a human, white blood cells attack the sensor as a foreign body, and, in the initial immune system onslaught, neutrophils are the primary white blood cells attacking the sensor. The defense mechanism of neutrophils includes the release of highly caustic substances known as reactive oxygen species. The reactive oxygen species include hydrogen peroxide.
In some non-limiting embodiments, the indicator 14 may be covered by a thin layer (e.g., 10 nm) on the outside of the indicator. The thin layer may protect against indicator molecule degradation. The thin layer may be platinum, and the platinum may be sputtered onto the outside surface of the indicator. Platinum rapidly catalyzes the conversion of hydrogen peroxide into water and oxygen, which are harmless to the sensor. The rate of this reaction is much faster than the boronate oxidation; thus, the platinum provides some protection against oxidation by reactive oxygen species. Although platinum is the catalyst of the conversion of hydrogen peroxide into water and oxygen in some embodiments, in alternative embodiments, other catalysts of this reaction, such as, for example, palladium or catalase, may be used for the thin layer instead of or in addition to platinum.
Under certain circumstances, after a sensor is assembled, it may be washed and then air dried (e.g., using heat, ambient air, dry gases or gases of a controlled humidity). When heat dried or air/gas dried, however, the hydrogel tends to shrivel like a dry sponge. This shriveling can cause a sputtered protective catalyst (such as platinum) to crack as the sensor is drying, which may pose a problem when the sensor is rehydrated and implanted. Cracking of protective catalyst may leave areas of the sensor vulnerable to attack by the immune system.
In addition to the physical damage (surface morphology) to the indicator or platinum layer due to shrinkage and cracking that accompanies air drying, another challenge involves long hydration periods that are required prior to implanting an air-dried sensor—especially after a period of storage. For air dried sensors, the average hydration period, or time to reach equilibration of in vivo signal, ranges from 30 minutes to 7 days depending on the storage time of the hydrogel (FIG. 2). This inconsistent behavior makes it difficult to predict the hydration profile of the hydrogel and time to stable baseline signal once inserted into a patient. Also, longer storage times in the dried state may cause irreversible signal loss that can never be regained due to the instability of the hydrogel in its unhydrated state (FIG. 2).
In the example shown in FIG. 2, the expected rehydration point of the hydrogel is shown to reach approximately 6 on the normalized signal scale. As can be seen in the plot, the signals of sensors that have been stored for 4 weeks reach the expected normalized signal level within a few days. Sensors stored for 8 weeks eventually reach the expected normalized signal level of 6, but it takes 5-6 days to reach that level. On the other hand, the signal level of sensor stored for 4 months or more never reach the expected normalized signal level.
Thus, as shown in FIG. 2, as the hydrogel is stored for periods greater than one month, the hydration of the hydrogel may take days or even months to equilibrate. Storage longer than 4 months does not allow the hydrogel to rehydrate at all.
Thus, a need exists for techniques for improving the longevity (shelf life) and rehydration characteristics of hydrogels employed in analyte sensors while minimizing or eliminating the surface morphology associated with air drying.